The present invention relates to superconducting magnets. It finds particular application in conjunction with superconducting magnets used in magnetic resonance imaging apparatus and will be described with particular reference thereto. It is to be appreciated, however, that the invention will also find application in other low temperature magnets in applications in which eddy magnetic fields are undesirable, such as nuclear magnetic resonance spectrometers and the like.
Magnetic resonance imaging systems superpose three magnetic fields over an imaging volume. A superconducting magnet is commonly used to provide a spatially large as well as a temporally and spatially constant main magnetic field. This main field is referred to as the B.sub.O field and is usually between 0.5 to 2.0 Tesla within an imaging volume of 50 cm diameter. A second magnetic field changes with time and has spatial linear .differential.B.sub.z /.differential.x, .differential.B.sub.z /.differential.y, .differential.B.sub.z /.differential.z gradients for spatial encoding that are aligned with the B.sub.O field. These are typically in the range of +/- 10 mT/m. Transverse to the direction of these two fields is an RF magnetic field used for magnetization reorientation during the NMR experiment.
In a conventional MRI magnet system, a set of B.sub.O field generating coils are configured in such a way that they provide only about a 10 micro Tesla variation within the 50 cm diameter imaging sphere. The coils are typically superconductors operating at a temperature near that of liquid helium, 4.2.degree. K. Typically, between six and eleven superconducting coils are series connected and immersed in a common liquid helium reservoir. In order to minimize the heat gain of the magnet from the room temperature environment, a cryostat is provided. The cryostat is an evacuated container to eliminate convective heat transfer. The magnet is immersed in a liquid helium bath at 4.2.degree. K. and near atmospheric pressure. Between these two containers two cold shields are suspended. One is cooled to about 10.degree.-20.degree. K. by refrigeration and/or the boil off helium gas. The other is typically cooled to about 77.degree. K. Several layers of aluminized mylar minimize the radiative heat transfer. The 77.degree. K. cold shield is either maintained by the boil off of liquid nitrogen or by mechanical refrigeration which can cool it to 60.degree.-70.degree. K. There is also a conductive heat transfer in the suspension members of these internal structures. This is, of course, minimized by design. The cryostat design is thus optimized to provide for a minimum heat gain and thus a minimum helium boil off.
A conventional MRI system surrounds the patient or part of the patient with an RF coil, then the gradient field coil, and finally the room temperature bore of the magnet. Although the cryostat does in general provide an excellent barrier to heat transfer, it also provides electrically conductive materials to support eddy currents from the time changing gradient fields. The magnetic fields resulting from these eddy currents may be substantial and require correction. Conventionally, this is done electrically with pre-emphasis. However, the most effective method is not to create the eddy currents in the first place rather than correcting for them after they are created. This is done with self shielded gradients which consist of a primary set of coils at a first radius as well as a secondary or shield set of coils at a second radius. The design of these coils is such as to minimize the eddy currents at the position of the magnet cold shields as described in U.S. Pat. No. 4,896,129 to Turner, et al. The efficiency of the self shield gradient is strongly influenced by the ratio of the shield coil and primary coil radii, because one partially cancels the field of the other over the imaging volume, in addition to cancelling fields outside of the coils.
There are three general types of superconducting magnets, unshielded, passively shielded, and actively shielded. By design, an unshielded magnet makes no provision to control the field external to the magnet. A passively shielded magnet is designed with iron close to the magnet so as to reduce the spatial extend of the magnet's field. An actively shielded magnet is designed with inner and outer coils (like the self shielded gradient) to minimize the magnetic field outside of the magnet. A hybrid magnet may use some active and some passive shielding. All of these types of magnets require a shimset in order to reach the desired level of homogeneity in the patient volume. A set of Garrett coils can be made to provide the additions to the main field to make the main field more uniform. Some of the coils of this shimset, either superconductive or resistive, can couple to the gradient coil, as does the shield coil of the self shielded gradient. To eliminate these interactions and reduce cost, most MRI systems now have a passive (iron) only shimset. This shimset is located in the room temperature bore of the magnet in the vicinity of the gradient coil.
More specifically, the first or higher temperature cold shields which are traditionally disposed along the patient bore and outside diameters of the vacuum chamber are cooled with liquid nitrogen to about 77.degree. K. For example, a cylindrical liquid nitrogen reservoir is provided inside the vacuum chamber around the outside of the superconducting magnet. The reservoir is thermally connected with a cylinder around the bore and disks at the edges, such that the nitrogen reservoir is cooled by conduction. However, this liquid nitrogen cooling is commonly replaced with mechanical refrigeration. The mechanical refrigeration unit is thermally connected with a copper or aluminum cylinder disposed adjacent the interior diameter of the vacuum chamber and another aluminum or copper cylinder disposed adjacent the outer diameter of the vacuum chamber. The mechanical refrigeration cools these cylinders to a temperature around 60.degree.-70.degree. K., the two cylinders being thermally linked at each end by disks.
To reduce the heat transport, second or inner cold shields are commonly disposed between the first cold shields and the liquid helium vessel. The second cold shields includes one copper or aluminum cylinder disposed between the inner diameter of the superconducting magnets and the inner cylinder of the first cold shield and a second copper or aluminum cylinder disposed between the outer diameter of the superconducting magnets and the outer cylinder of the first cold shield. The cylinder is thermally connected by disks at opposite ends. The second cold shield is thermally connected, such as by thermally conductive metal straps, with a second stage of the refrigeration unit which cools it to about 20.degree. K. Again, the cylinders are surrounded inside and out by layers of the super insulation.
A gradient field coil is disposed in the bore displaced from the inner wall of the cryostat for selectively generating magnetic field gradients across an imaging region in the bore. Typically, the gradient magnetic fields are applied for short durations, sufficiently short that they might more aptly be referenced as magnetic field gradient pulses. The gradient magnetic field pulses, particularly the rising and falling edges of the gradient field pulses, induce voltages in surrounding electrically conductive structures which, in turn, cause eddy currents. That is, eddy currents are generated in the magnet formers, the shimming coils, the liquid helium reservoir, the cold shields, and the like. At the very low resistance found at these temperatures, the eddy currents can persist for periods in excess of the repeat time of the magnetic resonance sequence. The generated eddy currents are very complex, varying with frequency, temperature, thickness of the electrically conductive structure, and the like. These generated eddy currents, in turn, generate eddy magnetic fields in the imaging region in the bore of the magnet. These eddy magnetic fields disrupt the precise magnetic fields that make for good quality, high resolution magnetic resonance imaging.
One technique for compensating for the eddy magnetic fields is through pre-emphasis. During initial pre-emphasis calibration, the contribution from the eddy magnetic fields are determined and the current pulses for generating the gradient magnetic fields adjusted accordingly. More specifically, the currents used to generate the gradient magnetic fields are adjusted such that the generated gradient field and the eddy magnetic field sum to produce the desired magnetic field in the image region. See for example, U.S. Pat. No. 4,761,612 issued Aug. 2, 1988 to Holland and Stauber entitled "Programmable Eddy Current Correction", and U.S. Pat. No. 4,703,275 issued Oct. 27, 1987 to Holland entitled "Method and Apparatus to Compensate Eddy Currents in Magnetic Resonance Imaging". Although pre-emphasis is effective, it does not completely correct for the complex gradient magnetic fields. The generated eddy magnetic field is very complex and cannot always be completely compensated for with a linear gradient.
One technique for reducing the eddy magnetic fields is to install a self-shielded gradient coil within the warm bore of the superconducting magnet. The self-shielded gradient coils include a gradient coil and a surrounding active shielding coil. The shielding coils are designed to cancel the gradient magnetic field at the positions of the cold shields to prevent inducing eddy currents, hence eddy magnetic fields. However, the shield coils generally require a diameter that is about 1.3 times the diameter of the primary gradient coil. This either reduces the size of the usable bore within the gradient coil or requires a larger, more expensive superconducting magnet. For medical diagnostic imaging, the minimum size of the bore is generally dictated by the dimensions of the human torso, hence reducing the useful bore size is undesirable.
The present invention contemplates a new and improved gradient shielding coil which overcomes the above-referenced problems and others.